Magnetic resonance signal detector grid assemblies for magnetic resonance imaging

ABSTRACT

A magnetic resonance (MR) signal detector grid assembly adapted for detection of signals in a magnetic resonance imaging apparatus is disclosed. The MR signal detector grid assembly may include a non-resonant grid of a plurality of wire elements and a plurality of detector elements. At least one end of a selected wire element may be electrically attached to an end of at least one other selected wire element. At least one of the detector elements may be configured to detect an MR signal induced on a wire element in the non-resonant grid. The plurality of wire elements may be substantially in the same plane, and the plane may be warped into an arbitrary three-dimensional surface. The three-dimensional surface may have a shape that is cylindrical, partially cylindrical, and/or saddle-shaped. The plurality of wire elements may form a rectilinear grid, a hexagonal grid, a triangular grid, and/or a pseudo-random grid.

CROSS-REFERENCE TO RELATED APPLICATIONS

The present application is a Continuation in Part of U.S. applicationSer. No. 14/814,002, filed Jul. 30, 2015, which claims priority to andthe benefit of U.S. Provisional Application Ser. No. 62/031,119, filedJul. 30, 2014, the disclosures of which are incorporated herein byreference.

BACKGROUND

In magnetic resonance imaging (MRI), surface coils are placed close tothe desired object to be imaged to maximize the signal-to-noise ratio(SNR). High performance surface coils are typically combined to formphased arrays comprised of overlapped single loop coils. These phasedarrays offer the SNR of small surface coils over a field-of-view (FOV)of a larger surface coil. Phased arrays are used to achieve highresolution imaging of different parts of the body. In addition tooffering a SNR advantage, multiple independent receive elements canenable accelerated imaging techniques.

Traditional phased array designs are limited because they have to becustomized to each body part. Customization criteria include shape andimpedance matching. Further, a large number of components are requiredfor this traditional approach.

SUMMARY

The present disclosure describes a magnetic resonance (MR) detector thatmay replace conventional MR imaging coil arrays. The present disclosuredescribes an example coil design approach that may reduce the number ofcomponents for MR imaging devices and may eliminate the need of tissuematching. This new approach implements a non-resonant grid in whichMR-induced currents are allowed to flow unconstrained over the grid(unlike conventional phased array coils in which current is constrainedto flow within each loop). Current in each element of the grid may bedetected with inductively-coupled pickup loops, which may be attached toindependent receiver channels of the MR imaging system. In one example,individual integrated balun pickup coils may be inductively coupled toeach grid element. Other connection arrangements, however, may beemployed if desired.

In a first aspect, the present disclosure discloses an example coilarray assembly adapted for detection of signals in a magnetic resonanceimaging (MRI) apparatus. The coil array assembly may include anon-resonant grid of transformer elements and pickup coils. Each pickupcoil may be inductively coupled to a corresponding transformer elementin the non-resonant grid. Each pickup coil may be configured to detectan MR signal induced on a corresponding wire element in the non-resonantgrid of wire elements.

In a second aspect, the present disclosure discloses a radio frequency(RF) coil assembly for use in a magnetic resonance imaging (MRI) system.The RF coil assembly may include coil array elements adapted to connectto a corresponding plurality of signal lines in the MRI system. Each ofthe coil array elements may include a non-resonant grid of transformerelements and pickup coils. Each pickup coil may be inductively coupledto a corresponding transformer element in the non-resonant grid. Eachpickup coil may include a loop coil body of conductor material and amatching circuit electrically coupled to the loop coil body and adaptedto be electrically coupled to one of the plurality of signal lines inthe MRI system. The loop coil body structure may include an integratedbalun.

In a third aspect, the present disclosure discloses a magnetic resonance(MR) signal detector grid assembly adapted for the detection of signalsin a magnetic resonance imaging (MRI) apparatus. The MR signal detectorgrid assembly may comprise a non-resonant grid of a plurality of wireelements and a plurality of detector elements. At least one end of aselected wire element may be electrically attached to an end of at leastone other selected wire element. At least one of the detector elementsmay be configured to detect an MR signal induced on a wire element inthe non-resonant grid. The plurality of wire elements may besubstantially in the same plane, and the plane may be warped into anarbitrary three-dimensional surface. The three-dimensional surface mayhave a shape that is cylindrical, partially cylindrical, and/orsaddle-shaped. The plurality of wire elements may form a rectilineargrid, a hexagonal grid, a triangular grid, and/or a pseudo-random grid.

BRIEF DESCRIPTION OF THE DRAWINGS

The foregoing and other features of the present disclosure will becomemore fully apparent from the following description, taken in conjunctionwith the accompanying drawings. Understanding that these drawings depictonly several embodiments in accordance with the disclosure and aretherefore not to be considered limiting of its scope, the disclosurewill be described with additional specificity and detail through use ofthe accompanying drawings.

In the drawings:

FIG. 1 depicts an exemplary MRI system in or for which the presentdisclosure may be implemented;

FIG. 2 depicts an example grid coil;

FIG. 3 depicts example transformer elements;

FIG. 4 depicts an example grid including example transformer elements;

FIG. 5 is a schematic representation of an example integrated balunpickup coil;

FIG. 6 depicts an example integrated balun pickup coil;

FIG. 7 depicts an example integrated balun pickup coil and example gridcoil; and

FIGS. 8-11 depict observations between example grid elements; eacharranged in accordance with at least some of the embodiments disclosedin the present disclosure.

DETAILED DESCRIPTION

The present disclosure describes a magnetic resonance (MR) detector thatmay replace conventional MR imaging coil arrays. The present disclosuredescribes an example coil design approach that may reduce the number ofcomponents for MR imaging devices and may eliminate the need of tissuematching. This new approach implements a non-resonant grid in whichMR-induced currents are allowed to flow unconstrained over the grid(unlike conventional phased array coils in which current is constrainedto flow within each loop). Current in each element of the grid may bedetected with inductively-coupled pickup loops, which may be attached toindependent receiver channels of the MR imaging system. In one example,individual integrated balun pickup coils may be inductively coupled toeach grid element. Each pickup coil may be configured to detect an MRsignal induced on a corresponding wire element in the non-resonant gridof wire elements. Other connection arrangements, however, may beemployed if desired. A pickup coil may also be referred to herein as a“detector element.”

MR imaging of internal body tissues may be used for numerous medicalprocedures, including diagnosis and surgery. In general terms, MRimaging starts by placing a subject in a relatively uniform, staticmagnetic field. The static magnetic field causes hydrogen nuclei spinsto align and precess about the general direction of the magnetic field.Radio frequency (RF) magnetic field pulses are then superimposed on thestatic magnetic field to cause some of the aligned spins to alternatebetween a temporary high-energy nonaligned state and the aligned state,thereby inducing an RF response signal, called the MR echo or MRresponse signal. It is known that different tissues in the subjectproduce different MR response signals, and this property can be used tocreate contrast in an MR image. An RF receiver detects the duration,strength, and source location of the MR response signals, and such dataare then processed to generate tomographic or three-dimensional images.

FIG. 1 shows an exemplary MRI system 100 in or for which MR imaging inaccordance with examples of the present disclosure may be implemented.The illustrated MRI system 100 includes an MRI magnet assembly 102.Since the components and operation of the MRI scanner are well-known inthe art, only some basic components helpful in the understanding ofsystem 100 and its operation will be described herein.

The MRI magnet assembly 102 may include a cylindrical superconductingmagnet 104 which generates a static magnetic field within a bore 105 ofthe superconducting magnet 104. The superconducting magnet 104 generatesa substantially homogeneous magnetic field within an imaging region 116inside the magnet bore 105. The superconducting magnet 104 may beenclosed in a magnet housing 106. A support table 108, upon which apatient 110 lies, may be disposed within the magnet bore 105. A regionof interest 118 within the patient 110 may be identified and positionedwithin the imaging region 116 of the MRI magnet assembly 102.

A set of cylindrical magnetic field gradient coils 112 may also beprovided within the magnet bore 105. The gradient coils 112 may alsosurround the patient 110 (or may surround the part of the patient's bodyof interest, such as the patient's hand and fingers). The gradient coils112 may generate magnetic field gradients of predetermined magnitudes,at predetermined times, and in three mutually orthogonal directionswithin the magnet bore 105. With the magnetic field gradients, differentspatial locations can be associated with different precessionfrequencies, thereby giving an MR image its spatial resolution. An RFtransmitter coil 114 surrounds the imaging region 116 and the region ofinterest 118. The RF transmitter coil 114 emits RF energy in the form ofa rotating magnetic field into the imaging region 116, including intothe region of interest 118.

The RF transmitter coil 114 may also receive MR response signals emittedfrom the region of interest 118. The MR response signals may beamplified, conditioned and digitized into raw data using an imageprocessing system 120, as is known by those of ordinary skill in theart. The image processing system 120 may further process the raw datausing known computational methods, including fast Fourier transform(FFT), into an array of image data. The image data may then be displayedon a monitor 122, such as a computer CRT display, LCD display, or othersuitable display.

In some examples, the RF transmitter coil 114 may include one or moregrid coils, such as grid coil 200 depicted in FIG. 2. Grid coil 200depicts an example 31-channel MRI grid coil design. FIG. 2 depicts a 31channel grid coil 200 design showing example positions of eachtransformer and integrated balun coil, the positions identified as 201,202, 203, 204, 205, 206, 207, 208, 209, 210, 211, 212, 213, 214, 215,216, 217, 218, 219, 220, 221, 222, 223, 224, 225, 226, 227, 228, 229,230, and 231. FIG. 2 also depicts two dashed boxes that represent thesize and shape of the large loop coil 240 and small loop coil 250 thatwere built to compare the performance of the grid coil 200, which willbe discussed herein.

A rectangular grid in accordance with the grid coil 200 design may beconstructed of wire elements 300 including a transformer loop using wiresuch as 12 AWG coated copper, for example. The wire elements 300 may betwisted or otherwise formed in a loop configuration such asconfigurations 360, 365 depicted in FIG. 3. The wire elements 300 mayeach include leg segments and a loop segment. For example, wire element300 having configuration 360 may be formed so as to have two legsegments 310, 315 and a loop segment 320. Similarly, wire element 300having configuration 365 may be formed so as to have two leg segments330, 335 and a loop segment 340. The wire elements 300 shown in FIG. 3may be soldered together or otherwise fused or formed together or withother wire elements. For example, the leg segment 310 may be soldered toleg segment 330.

An example grid 400 in which wire elements 401-431 were solderedtogether is depicted in FIG. 4. This example grid 400 is on a 12″×9″ FR4composite board, which includes wire elements 401-431 fastened to theFR4 composite board. Example grid 400 was constructed and tested toinvestigate the benefits of the grid coil 200 design. In this manner,wire elements having transformer loops were positioned at the positionsidentified as 401, 402, 403, 404, 405, 406, 407, 408, 409, 410, 411,412, 413, 414, 415, 416, 417, 418, 419, 420, 421, 422, 423, 424, 425,426, 427, 428, 429, 430, and 431, and all of the wire elements 401-431were soldered together. In other words, for at least a portion of thewire elements 401-431, each end of a selected wire element may beelectrically attached or connected to an end of at least one otherselected wire element.

It should be noted that the present disclosure is not limited to arectilinear grid. Other configurations such as rectangular grids,triangular grids, hexagonal grids, and arbitrarily connected networksmay be desirable. In particular, it may be useful to let each wireelement in the grid have a unique length and orientation with respect toother wire elements in such a fashion that optimizes MR signalsensitivity for selected regions of interest. Furthermore, the grid neednot be restricted to a plane. It may be advantageous to deform the gridinto a cylinder, curved surface (e.g. a saddle shape), or arbitrarilyshaped surface intended to closely conform to a selected region ofanatomy.

The wire elements may be arranged substantially in the same plane. Theplane may be warped into an arbitrary three-dimensional surface. In oneexample, the three-dimensional surface may be a cylinder orsubstantially a cylinder. In another example, the three-dimensionalsurface may be a portion of a cylinder or substantially a portion of acylinder. In yet another example, the three-dimensional surface may be aportion of a saddle shape or substantially a portion of a saddle shape.The saddle shape may have both positive and negative curvature.

The wire elements may additionally or alternatively have otherarrangements. In one example, the plurality of wire elements may form arectangular grid having a substantially uniform spacing between parallelwire elements. In another example, the plurality of wire elements mayform a hexagonal grid having a substantially uniform spacing betweenparallel wire elements. In another example, the plurality of wireelements may form a triangular grid having a substantially uniformspacing between parallel wire elements. In another example, theplurality of wire elements may form a pseudo-random grid wherein eachwire element has a selected length and angular orientation that ispotentially different to other wire elements. The pseudo-random grid maybe configured to optimally detect MR signals from an object having aselected size and geometry.

A signal detection mechanism may rely on inductive coupling between aselected one of the wire elements and a selected one of the detectorelements. However, it should be noted that the present disclosure is notlimited to inductive coupling between the wire elements and the MRIsystem. Other coupling mechanisms such as capacitive and opticalcoupling may be advantageous. In particular, it may be advantageous toplace a preamplifier directly on each wire element and connect theoutput of each preamplifier to the MRI system via a coaxial cable.Likewise, it may be advantageous to place both a preamplifier and Analogto Digital Converter (ADC) on each wire element to permit an opticalconnection between each wire element and the MRI system.

A substantial benefit of the present disclosure is the ability to assignunique weights to the MR signals detected by each wire element. Theseweights can be pre-calculated for each location within the imagingvolume of the MRI system. This allows each pixel in the image to bereconstructed with a unique set of wire element weights, which in effectpermits MR imaging using virtual coils. These virtual coils canrepresent coils which cannot be physically constructed and/or placed.For example, it is possible to create a virtual coil that would beinside the patient's body if it were physically constructed. Anotheradvantageous aspect of the present disclosure is that the weightsassigned to each wire element can provide quadrature sensitivity forevery point in the imaging volume. This provides a signal to noise ratioadvantage over traditional loop MR coil arrays which have limited or nosensitivity to quadrature MR signals. The MRI apparatus may use aselected mathematical weight for MR signals acquired from each wireelement. The weight may be computed for each voxel in a final magneticresonance image.

Integrated balun coils 500, 600 may be coupled to the wire elements viafasteners (e.g., nylon nuts, bolts, and washers). A circuit schematic ofan example integrated balun coil 500 is shown in FIG. 5. FIG. 6 depictsan example integrated balun coil 600 arranged in accordance with thecircuit schematic of FIG. 5. More details regarding the design ofexample integrated balun coils may be found in U.S. patent applicationSer. No. 14/627,680, filed Feb. 20, 2015; the disclosure of which isincorporated herein by reference.

Integrated balun coil 500 may include a first copper tube segment 514(or wire, trace or other sort of conductor segment) and a second coppertube 510 (or wire, trace or other sort of conductor) mounted to amatching network circuit 505. The bottom sections of the first andsecond copper tube segments 514, 510 are connected to a copper tube stemsegment 516 extending generally axially (generally perpendicularly) fromthe first and second copper tube segments 514, 510 and is mounted to thematching network circuit 505 board such that the circuit board alsoextends generally perpendicularly from the loop coil 500. An innerconductor 524 and a dielectric insulator 525 extend through the coppertube segments 514, 516 forming a coaxial cable 530 in such segments 514,516, where the dielectric insulator 525 separates the inner conductor524 from the interior of the copper tube segments 514, 516. The coaxialcable 530 extends from the matching network circuit board 595 throughthe stem segment 516 and up through the first copper tube segment 514.At the top of the copper tube segments 514, 510 a tuning capacitor 520is electrically coupled between the second copper tube segment 510 tothe inner conductor 524 of the coaxial cable 530. The inner conductor524 of the coaxial cable 530 is also coupled to the matching network 505provided on the circuit board.

As depicted in this diagram, the inner conductor 524 of the coaxialcable 530 is coupled between the matching circuit 505 and the tuningcapacitor 520. On the other hand, the outer conductor (the copper tubematerial of the first segment 514 and the stem segment 516) of thecoaxial cable 530 is coupled to the matching circuit 505 at the stemsegment 516 end and is unconnected (open 512) at the other end.

The matching network 505 may include a pair of matching capacitors 540,545 and a matching inductor 550. The matching network 505 can berealized as impedance controlled microstrip line, capacitors orinductors, impedance controlled coaxial transmission lines, or acombination of them as will be apparent to those of ordinary skill. Thecircuit may also include a PIN diode 560 coupled between the first andsecond circuit leads 570, 580.

FIG. 6 depicts an example integrated balun coil 600 arranged inaccordance with the circuit schematic of FIG. 5. As shown in FIG. 6, anexemplary embodiment of a loop coil 600 according to the presentdisclosure includes a pair of segments: a first arcuate copper tubesegment 644 and a second arcuate copper tube segment 642 facing eachother to form a partially enclosed loop. The bottom sections of thefirst and second copper tube segments 642 and 644 are connected to acopper tube stem segment 646 extending generally axially (generallyperpendicularly) from the first and second copper tube segments 644, 642and is mounted to the matching network circuit board 650 such that thecircuit board 650 also extends generally perpendicularly from the loopcoil 600. An inner conductor 654 and a dielectric insulator 655 extendthrough the copper tube segments 644 and 646 forming a coaxial cable 648in such segments 644, 646, where the dielectric insulator 655 separatesthe inner conductor 654 from the interior of the copper tube segments644 and 646. The coaxial cable 648 extends from the matching networkcircuit board 650 through the stem segment 646 and up through the firstcopper tube segment 644. At the top of the copper tube segments 642 and644 a tuning capacitor 652 is electrically coupled between the secondcopper tube segment 642 to the inner conductor 654 of the coaxial cable648. The inner conductor 654 of the coaxial cable 648 is also coupled tothe matching network 651 provided on the circuit board 650.

FIG. 7 depicts an example assembly 700 coupling an integrated balun coilassembly to a grid coil wire element having leg segments 710, 715 and aloop segment 720. The integrated balun coil assembly in this exampleincludes an integrated balun coil loop 705, a circuit board 750, and anon-conductive base 760 for the integrated balun coil assembly. may be aLexan block or other non-conductive materials.

Integrated balun coil loop 705 may be coupled to each grid coil loop 720on the grid using a non-conductive alignment cradle 730. In someexamples, the non-conductive alignment cradle 730 may include nylonnuts, nylon bolts, and/or nylon washers. Other non-conductive materialsmay also be used to construct the non-conductive alignment cradle 730.The non-conductive alignment cradle 730 secures the integrated baluncoil loop 705 and the loop segment 720 of a grid coil coaxially acrossfrom each other. The non-conductive alignment cradle 730 also maintainspositions of the integrated balun coil loop 705 and the loop segment 720of a grid coil relative to each other. The non-conductive base 760 forthe integrated balun coil assembly may be a Lexan block or othernon-conductive material.

An example 31 channel grid coil with integrated balun coil configurationwas constructed in accordance with the configuration shown in FIG. 4.The dimensions of the squares were chosen to be 70 mm×70 mm afterempirically determining the distances required for −30 dB couplingbetween pickup coils. Each of the 31 transformer elements wereindividually cut and bent out of 12 gauge coated copper wire as depictedin FIG. 3. They were then soldered into a 3×4 rectangular grid andmounted on a FR4 12×9 inch board. Twenty-eight 300 pF capacitors weresoldered at selected joints to interrupt eddy currents. The grid wasmodified from the symmetric square pattern to a rectangular pattern tominimize use of capacitors, measuring 220 mm×285 mm in total.

Example integrated balun coils were constructed using 2.196 mm diametersemi-rigid coaxial cable. The diameter of the coils matched the diameterof the transformer elements on the grid for optimal inductive coupling.Each example pickup coil is composed of a segment of the copper coaxialcable, a small feeder board, multiple capacitors for tuning, and a diodeand an inductor for detuning during transmit. Each pickup coil wasindividually tuned to 127.74 MHz, the Larmor frequency of protons at 3T.

MR images 800 were taken with the example 31 channel grid coil on aPhilips Achieva 3 Tesla MRI system and were compared to images 900,1000, 1100 taken with a size-matched large loop coil 240 (see images1100), a small loop coil 250 (see images 1000), and a commercially-made16 channel array (see images 900). These images are shown in FIGS. 8-11.

Network measurements were performed with a Rhode & Schwarz ZNC 3 NetworkAnalyzer. The coupling of the transformer to the pickup coil was foundto be −3.2 dB. Isolation between individual grid elements was alsomeasured and found to be better than typically found in conventionalphased array coils.

A large rectangular traditional coil (the size of box 240) measuring 215mm×295 mm and a small square traditional coil (the size of box 250)measuring 75 mm×75 mm were constructed. Images were acquired with theexample 31 channel grid coil and the two traditional coils 240, 250 tocompare their sensitivity profile and SNR.

MR images 800, 900, 1000, 1100 were acquired with a Philips 3 T Achieva(Philips Healthcare, Best, Netherlands) MRI system. For phantom imaging,a T2 weighted turbo spin echo sequence was used with a 90° flip angle,TR=300 msec, a TE=4.5 msec, an FOV of 400 mm×400 mm, and a slicethickness of 4 mm. Four 2-liter bottles were used as phantoms, eachcontaining a solution of 2 g/l NaCl₂ and 1 g/l CuSO₄. These phantomsprovided a large homogenous medium for the evaluation of sensitivityprofile of the imaging coils.

In this example, the highest inter-channel coupling was observed at thepositions between adjacent grid element positions 411 and 412. Thiscoupling was observed to be −13.3 dB. 94.6% of coil couplings had betterthan −20 dB isolation.

Signal to noise measurements were taken both close to the coil where thesignal was at its highest intensity and in the center of the phantoms.Close to the surface of the phantom the grid coil (depicted in FIG. 8)had the highest SNR followed by the small loop (depicted in FIG. 10),large loop (depicted in FIG. 11), and 16 channel commercial phased array(depicted in FIG. 9) with signal-to-noise ratios of 36.08, 35.35, 14.48,and 7.54 respectively. In the center of the phantoms, the grid coil hadthe second highest SNR (12.48) with the small loop having an SNR of20.67

In these examples, the grid coil had better SNR than the two traditionalcoils 240, 250 tested and required fewer parts to construct than thecommercial phased array. Furthermore, the grid coil showed betterisolation between channels than the traditional phased array. Thisfeature alone may provide enhanced image performance for highacceleration factors. Further, the grid approach may use currentmeasurements from individual grid elements to synthesize other coils,including coils that cannot be physically constructed. (e.g. ahypothetical surface coil disposed inside the body with itscorresponding SNR advantage for small regions of interest such as asingle chamber of the heart). The grid approach may be furtheroptimized, and new grid layouts may include other shapes and/orconfigurations, including offset squares, hexagons, and octagons, forexample. This may maximize or improve SNR, image acceleration factors,and/or image uniformity.

Some examples of the present disclosure may eliminate the need forspecialized surface coils. In some examples,

-   -   Synthetic coils may be created with a single grid. This may        eliminate the need for specifically designed surface coils.    -   The grid does not load to the body. This may make the coil        easier to construct.    -   High acceleration factors are likely due to the increased        isolation between imaging channels.    -   The grid approach improves SNR with multiple channels while        maintaining a large FOV.    -   Every pickup coil in the grid may be identical. This may improve        manufacturability and robustness.

While the foregoing disclosure includes many details and specificities,it is to be understood that these have been included for purposes ofexplanation and example only, and are not to be interpreted aslimitations of the present disclosure. It will be apparent to thoseskilled in the art that other modifications to the embodiments describedabove can be made without departing from the spirit and scope of thepresent disclosure. Accordingly, such modifications are to be consideredwithin the scope of the present disclosure.

Likewise, it is to be understood that it is not necessary to meet any orall of the identified advantages or objects of the present disclosure inorder to fall within the scope of the present disclosure, since inherentand/or unforeseen advantages of such inventions may exist even thoughthey may not have been explicitly discussed herein.

All publications, articles, patents and patent applications cited hereinare incorporated into the present disclosure by reference to the sameextent as if each individual publication, article, patent application,or patent was specifically and individually indicated to be incorporatedby reference.

What is claimed is:
 1. A magnetic resonance (MR) signal detector gridassembly adapted for detection of signals in a magnetic resonanceimaging (MRI) apparatus, comprising: a non-resonant grid of a pluralityof wire elements, wherein at least one end of a selected wire element iselectrically attached to an end of at least one other selected wireelement; and a plurality of detector elements, each detector elementbeing configured to detect an MR signal induced on a given one of thewire elements in the non-resonant grid.
 2. The MR signal detector gridassembly of claim 1, wherein the plurality of wire elements aresubstantially in the same plane.
 3. The MR signal detector grid assemblyof claim 2, wherein the plane is warped into an arbitrarythree-dimensional surface.
 4. The MR signal detector grid assembly ofclaim 3, wherein the three-dimensional surface is substantially acylinder.
 5. The MR signal detector grid assembly of claim 3, whereinthe three-dimensional surface is substantially a portion of a cylinder.6. The MR signal detector grid assembly of claim 3, wherein thethree-dimensional surface is substantially a portion of a saddle shape.7. The MR signal detector grid assembly of claim 1, wherein theplurality of wire elements form a rectilinear grid having asubstantially uniform spacing between parallel wire elements.
 8. The MRsignal detector grid assembly of claim 1, wherein the plurality of wireelements form a hexagonal grid having a substantially uniform spacingbetween parallel wire elements.
 9. The MR signal detector grid assemblyof claim 1, wherein the plurality of wire elements form a triangulargrid having a substantially uniform spacing between parallel wireelements.
 10. The MR signal detector grid assembly of claim 1, whereinthe plurality of wire elements form pseudo-random grid.
 11. The MRsignal detector grid assembly of claim 10, wherein each wire element hasa selected length and angular orientation that is different than atleast one of the other wire elements.
 12. The MR signal detector gridassembly of claim 10, wherein the pseudo-random grid is configured tooptimally detect MR signals from an object having a selected size andgeometry.
 13. The MR signal detector grid assembly of claim 1, whereinsaid signal detection relies on inductive coupling between a selectedone of the wire elements and a selected one of the detector elements.14. The MR signal detector grid assembly of claim 1, wherein said signaldetection relies on capacitive coupling between a selected one of thewire elements and a selected one of the detector elements.
 15. The MRsignal detector grid assembly of claim 1, further comprising apreamplifier connected to a selected one of the wire elements.
 16. TheMR signal detector grid assembly of claim 1, wherein one or moredetected MR signals are transmitted to the MRI apparatus via a coaxialcable.
 17. The MR signal detector grid assembly of claim 1, wherein saidMRI apparatus uses a selected mathematical weight for MR signalsacquired from each wire element, said weight being computed for eachvoxel in a final magnetic resonance image.